1. Field of the Invention
The present invention is directed to a device, method, and system for measuring respiratory resistance of living organisms, in particular human beings. Respiratory resistance is a measurement which has significant clinical and physiological interest. Increased resistance can typically be related to an assortment of respiratory diseases such as asthma, bronchitis, pneumonia, emphysema, and various other obstructive disorders. Resistance measurements can be useful for evaluating the respiratory effects of bronchoconstrictive and bronchodilatory drugs, as well as airborne contaminants and natural particulates. Energy expenditure for respiratory functions can constitute a significant part of a body's total expenditure during exercise. Energy expenditure is increased when protective masks are worn, which is common in occupations wherein exposure to contaminated environments can occur. Respiratory resistance measurement, therefore, is an important measurement for many different purposes. The invention is directed to a respiratory resistance measuring device which is reasonably inexpensive, easy to use, and accurate.
2. Description of the Related Art
Several techniques are available to measure various aspects of respiratory resistance. Among those which are non-invasive, the technique of forced oscillation measures total respiratory resistance, but it requires a great deal of expertise and subject cooperation. Whole-body plethysmography measures the resistance of the airways alone, but it requires a large apparatus that is not portable or easy to use. Both techniques are quite expensive. Other techniques are invasive. A measurement of pressure that can be used to calculate part of respiratory resistance is made by the insertion of an esophageal catheter directly into the subject. These conventional measurement techniques are not appropriate for important groups of patients such as neonates, pre-school children, and the critically ill.
Airflow perturbation techniques were proposed by several groups in the 1970s as a more simple way of obtaining respiratory resistance. Dr. Arthur Johnson's interest in measuring airways resistance led him to design the first Airflow Perturbation Device (APD) in 1974, which is discussed in U.S. Pat. No. 4,220,161, which is hereby incorporated by reference. The APD was intended to be a simple, economical, non-invasive device for the measurement of airway resistance, as discussed in Johnson and Lin, Airflow Resistance of Conscious Boars, Transactions of the ASAE, volume 26, pages 1150-1152, 1983. This and other earlier work has produced measurements in the expected range for test subject airway resistance, but further conclusions have not been drawn. Airway resistance measurements are discussed in U.S. Pat. No. 4,856,532, the contents of which is hereby incorporated by reference. The present invention is directed to an improved Airflow Perturbation Device, as well as respiratory resistance measuring system and method. The APD of the present invention measures not airway resistance but respiratory resistance, is sensitive to resistance changes, and the measurement is correlated to airway resistance.
In order to properly understand the operation of an airflow perturbation device according to the invention, a brief background regarding respiratory mechanics is helpful.
The energy required by breathing is primarily the energy used for inspiration, or inhalation. Expiration, or exhalation, is usually a passive process driven by the elastic recoiling of the lungs and chest wall. The work of inspiration has three fractions. Work is done to overcome compliance, resistance, and inertance impedances. Compliance work expands the lungs and the enclosing chest cage against their elastic forces. Resistance work is done to overcome the viscous and frictional resistance to air flow and tissue movement. The measurement of this resistance is the primary goal of the APD. Inertance work accelerates the mass of the volume of air inhaled, the lung tissue and chest wall.
During normal quiet breathing, most of the work done is the compliance work needed to expand the lungs. A smaller amount is lost in overcoming tissue resistance. A similar amount of energy is lost to airways resistance. Mass inertia work is considered negligible, as shown in FIG. 1. During heavy breathing, such as during exercise, air must flow through respiratory passageways at a much higher velocity. The energy dissipated by airways resistance then becomes the greatest proportion of respiratory work.
These terms of compliance, resistance, and inertance are completely analogous to the electrical terms of capacitance, resistance, and inductance. This allows the use of electrical symbols in illustrating models of the respiratory system. FIG. 2 shows a respiratory model in terms of electrical components of the system made up of airways, lung tissue, and chest wall fractions.
A pressure balance equation of this model where the ground point is atmospheric pressure produces equation 1. EQU (p.sub.mo -p.sub.atm)+(p.sub.alv -p.sub.m)+(p.sub.pl -p.sub.alv)+(p.sub.mus -p.sub.pl)+(p.sub.atm -p.sub.mus)=0 (1)
In this formula, p.sub.mo is mouth pressure, p.sub.atm is atmospheric pressure, p.sub.alv is alveolar pressure, p.sub.pl is pleural pressure, p.sub.mus is muscle pressure. The mouth-atmosphere term is nonzero when wearing a mask, breathing through a tube, or breathing through any other obstruction. The APD functions by periodically increasing this term. Pleural pressure is at the lung-chest wall interface. Alveolar pressure is at the air-lung tissue interface. Muscle pressure is exerted by the diaphragm.
The pressure differences between mouth, alveolar, pleural and muscle levels can be expressed by the equation for a linear circuit. ##EQU1## where
______________________________________ .DELTA.P: Pressure Difference V: Volume C: Compliance V': Flowrate R: Resistance V": Volume acceleration I: Inertance ______________________________________
The complexity of FIG. 2 can be reduced to the model shown in FIG. 3. Resistance and inertance in this model are the sums of airways, lung tissue, and chest wall components. Compliance is shown as a function of lung tissue and chest wall compliances. Airways compliance is nearly zero, so it is neglected.
Rohrer's relationship has been generalized to include different lung volumes. ##EQU2## where
______________________________________ R: Airway resistance V: Lung volume K.sub.1 : First coefficient V': Flowrate K.sub.2 : Second coefficient V.sub.res : Lung residual volume K.sub.3 : Third coefficient ______________________________________
The last fifteen years has seen much work to document the characteristic frequency dependence of respiratory resistance. While airway resistance is not highly frequency dependent at low frequencies, total respiratory resistance is. In healthy humans, respiratory resistance is relatively constant from 3 to 10 Hz. It increases at lower and at higher frequencies. More discussion of frequency dependence will follow, along with the topic of forced oscillation of the respiratory system.
If inertance and compliance are negligible, resistance equals pressure divided by flowrate. Respiratory airflow is easy to measure. Almost any human subject can breathe through a flowmeter. Pressures inside the body are much more difficult to measure, complicating the measurement of resistance. Airway resistance is that between the mouth and the alveoli. Pulmonary resistance is airway resistance plus lung tissue resistance. Respiratory resistance is pulmonary resistance plus chest wall resistance. If inertance and compliance are considerable, pressure divided by flowrate is defined as impedance. The portion of impedance consisting of inertance plus compliance is termed reactance.
Five methods have been used to measure various aspects of respiratory resistance or impedance. They are the esophageal balloon, body plethysmograph, forced oscillation technique, flow interrupter, and APD. Their abilities and limitations vary.
The esophageal balloon method, which measures pulmonary resistance, is the most direct measurement of the five. A balloon-tipped catheter is inserted through the nose down to the lower third of the esophagus. The pressure at the catheter tip is assumed to equal pleural pressure, or the pressure just outside the lung tissue. Pressure difference is measured between the catheter tip and the subject's mouth. The directness of this measurement makes it useful for validating other methods of resistance measurement. Its invasive nature limits its clinical use.
The interrupter method was one of the earliest attempts to non-invasively measure airways resistance. Neergaard and Wirz first used airflow interruption in 1927. With this method, airflow is suddenly halted and mouth pressure is monitored. The mouth pressure measured immediately upon interruption is assumed to equal alveolar pressure just prior to the interruption. It is divided by the airflow rate just prior to the interruption to produce airway resistance. This assumption is not completely valid. It neglects the effects of air mass inertance, airways compliance, and frequency dependent lung tissue viscoelastic parameters. Theoretical analysis has also suggested that upper airways compliance may cause the interrupter to underestimate resistance. Experimental interrupter data has been obtained from open-chested dogs, or dogs who have had the breastplate surgically moved to allow direct measurement. It has shown the interrupter to measure airways resistance and not pulmonary resistance. Experimental data obtained with closed-chested dogs indicates that the measurement is airways plus chest wall resistance. A compact, portable interruption device suitable for clinical use has been developed and tested against the body plethysmograph in adult human subjects. It has been shown to produce measurements correlated to, but larger than, the plethysmograph. It has also measured resistance changes from bronchodilator treatment similar to those measured by plethysmography.
DuBois introduced whole-body plethysmography for measuring airways resistance in 1956. His constant-volume plethysmograph, or body box, is a sealed container in which the subject sits and breathes. The plethysmograph had been previously used to measure thoracic gas volume by application of Boyle's Law. During exhalation, alveolar pressure exceeds box pressure as the lungs expel air through the airways. Since the box is sealed, air compression inside the lungs results in lower air pressure throughout the remainder of the box. During inhalation, air expansion inside the lungs results in greater air pressure throughout the rest of the box. Box pressure, inversely proportional to lung pressure, is monitored by a sensitive pressure transducer. Lung pressure can be measured by occluding the airflow and measuring mouth pressure. In this static system, mouth pressure and alveolar pressure are the same. Therefore, the conversion factor between box pressure and alveolar pressure is the slope of the mouth pressure-box pressure curve.
The classic DuBois airways resistance test begins with shallow panting through a pneumotach flowmeter. The technician waits for air flowrates between -0.5 and +0.5 L/s. After several seconds of box pressure-airflow data collection, a shutter closes to occlude airflow and several seconds of box pressure-mouth pressure data is collected. The conversion factor from the occlusion data is applied to the box pressures in the panting data to produce what is believed to be an alveolar pressure-flowrate curve. The slope of that curve is plethysmographic airways resistance.
Sealing a human subject inside a plethysmograph poses a thermal problem. Body heat production heats and compresses the air, so boxes are built with slow leaks or ventilation valves to release the pressure. The error caused by slowly building pressure, a low frequency signal, is minimized by performing the test with panting, a high frequency signal. An alternative instrument less sensitive to this thermic error is the volume displacement plethysmograph. It works similarly to the constant-volume plethysmograph, except that changes in box volume are monitored instead of changes in box pressures. This less-common plethysmograph is more mechanically complex and suffers a more limited frequency response.
Another problem, recognized by DuBois, is that the air volume exhaled through the pneumotach can be different than the volume inhaled due to heating, humidification, and to a much lesser degree, respiratory gas exchange. Several solutions to the heating and humidification problem have been attempted. DuBois suggested that the shallow panting minimizes this problem because the inhaled air will be pneumotach dead space air, already warm and moist. A relationship between panting frequency and plethysmographic airways resistance has been found by numerous researchers and attributed to the noninstantaneous air heating and wetting. Lower frequency panting results in lower airway resistance values. For accurate measurements at frequencies below 2 Hz, the subject can breathe air already conditioned to body temperature and saturated. Some boxes are equipped with rubber bags to hold warm, saturated air, but cost and sanitary considerations limit their use. Another alternative to deal with this problem is to mathematically remove this effect. A body box has been marketed with "Electronic BTPS correction," but this technique has been found to be inaccurate.
Body plethysmographs have found use among scientists and clinicians. Among other things, they have been used to measure the pressure-volume-flow characteristics of the lungs, test the effectiveness of bronchodilators, respiratory responses to temperature and humidity, and bronchoconstriction due to smoking. Yet this technique suffers from high technician variability and poor reproducibility. In the case of high-resistance subjects, researchers have found consistent underestimations of airways resistance. The assumption that mouth pressure equals alveolar pressure during occlusion generally holds in healthy patients. In patients experiencing spontaneous or induced bronchoconstriction, however, significant mouth-alveolar flows cause a pressure gradient. This results in a high thoracic gas volume and low airways resistance measurement.
The forced oscillation technique (FOT) was also introduced by DuBois as a way to measure respiratory resistance, the sum of airways, lung tissue, and chest wall resistances. A piston pump was used to apply sinusoidal pressure oscillations at the mouth. The pressure oscillations were compared to the resulting flow oscillations to obtain magnitude (.vertline.Z.sub.rs .vertline.) and phase angle (.theta..sub.rs) of the respiratory system. Flow led pressure (.theta..sub.rs was negative) at low frequencies indicating a prominent compliance effect. Near 6 Hz, pressure and flow were in-phase and .vertline.Z.sub.rs .vertline. reached its minimum value where it represented respiratory resistance alone. From this, 6 Hz is considered the resonant frequency of the average respiratory system. Pressure led flow (.theta..sub.rs was positive) at higher frequencies indicating a prominent inertial effect. Oscillations were induced from 2 to 15 Hz. Impedances at resonant frequencies were found to range from 2 to 4 cmH.sub.2 O/L/s.
Since this initial work, many researchers have modified the FOT to make the calculations easier or the measurements more enlightening. Goldman et al. (1970) measured respiratory resistance at frequencies other than at the resonant frequency with an alternative technique. With sinusoidal forcing, there are two instants per cycle when flow acceleration is zero and volumes are identical. These are at the extremes of flow. Respiratory resistance was calculated as the change in pressure between these points divided by the change in flow rate. Resistances in six healthy subjects were found to range from 1.7 to 2.5 cmH.sub.2 O/L/s. While these early measurements with FOT were limited to discrete frequencies, some researchers felt that additional information about the respiratory system could be gained by oscillating with a wide range of frequencies simultaneously. The development of the fast Fourier transform (FFT) for microcomputers in the 1970's enabled FOT to measure impedance in the frequency domain. The FFT of mouth pressure is divided by the FFT of flowrate to produce impedance as a function of frequency. A polyfrequent forcing function, either designed or random, is applied with a loudspeaker or piston pump. Impedances are now typically reported as having a real component and an imaginary component. The real component represents the resistance, or the in-phase component. The imaginary component represents reactance, or the out-of-phase component. Reactance is the sum of inertance and compliance impedances. No consistent differences have been found between impedances derived from discrete frequency or from broadband input frequencies ranging from 4 to 256 Hz.
Researchers usually calculate FOT resistance at relatively high frequencies in order to avoid the low frequencies associated with normal breathing. Some measurements are also taken at low frequencies, but they require a higher degree of subject cooperation. Subjects must be trained to maintain voluntary apnea for periods of approximately 30 seconds.
In healthy humans, respiratory resistance decreases sharply from 0 to 2-3 Hz. This decrease occurs almost entirely within the chest wall, with pulmonary resistance remaining constant. Respiratory resistance remains somewhat stable or decreases slightly from 2-5 Hz up to 20 Hz. Measurements conducted from 4 to 256 Hz generally indicate that respiratory resistance increases with frequency, reaching a peak near 160 Hz and then falling off at higher frequencies. Very similar results have been found in dogs. Less data exists for the effect of flowrate and volume upon the tissue resistances. It has been shown that lung tissue resistance depends upon volume. Because of these, and possibly other dependencies, resistance values must always be presented accompanied by information regarding the conditions under which they were obtained.
As with the interrupter technique, the compliance of the upper airways such as the cheeks are a concern. The dampening of the induced pressure signal here is referred to as upper airway shunting. Habib and Jackson (1993) investigated ways of reducing shunting error. FOT was implemented by applying the signal in the conventional manner, through the mouth, with the cheeks supported by the hands and also without hand support. Additionally, FOT was implemented through the mouth while the head was submerged in a rigid water tank. FOT was also implemented by applying oscillations to a rigid chamber in which the head was sealed. Standard hand cheek support was found to most closely represent shunt-free impedance.
Airflow perturbation techniques for measuring resistance were independently introduced at least four times in the 1970's. These techniques mechanically resembled the airflow interrupter but were operationally similar to the Goldman et al. (1970) implementation of FOT. Sobol (1970) introduced a device consisting of a bifurcated tube with one unobstructed flowpath and one screened flowpath. A shutter directed air alternately through either flowpath. Airways resistance was calculated from the ratio of unscreened to screened flowrates during inhalation. Flowrates were measured with a hot-wire anemometer. The calculation also required screen resistance. Both healthy subjects and respiratory patients were tested with the device. A strong correlation between the new technique and plethysmographic resistance was found, with Sobol's resistances being consistently higher.
Kures (1974) introduced a perturbation device constructed from a pneumotach with an attached shutter that could be closed to increase resistance without completely occluding flow. His "Additive Technique" calculated resistance from the ratio of the mouth pressure perturbation magnitude to the flow perturbation magnitude. Unlike Sobol's method, which required the knowledge of screen resistance, this method did not require knowledge of the shutter resistance. Healthy adults, healthy children, and asthmatic children were tested. These resistances were reported as the average of inspiratory and expiratory values. They were compared to plethysmographic resistance and found to be both correlated and of the same magnitude.
At about the same time, Johnson et al. (1974) introduced the APD as disclosed in U.S. Pat. No. 4,220,161. This device has one flowpath with a screened resistance mounted on a rotating wheel. The screened wheel provided for smooth, quick switching from substantially unperturbed to maximally perturbed flow. Resistance was calculated using the mouth pressure perturbation magnitude relative to the known screen resistance. Later, the APD resistance calculation dispensed with the use of screen resistance and calculated airway resistance as the magnitude of the mouth pressure perturbation divided by the magnitude of the flowrate perturbation. Measurements taken from healthy subjects were found to be similar in magnitude to plethysmographic resistance. Measurements taken in live pig lungs with an intraluminal catheter have shown perturbations to extend to the twelfth airway bifurcation, but not necessarily beyond. APD resistances tended to be higher in the expiratory direction than in the inspiratory direction. They tended to increase with flowrate.
Schmid-Schoenbein and Fung (1978) built a perturbation device similar in appearance to the APD but with a partially-closing shutter in place of the rotating screen. They attempted to determine not just resistance, but inertance and compliance as well, by taking multiple samples during the perturbation. Their technique involved digitally sampling about 10 points over the course of a 30 to 80 ms perturbation. Perturbations were visually selected. The three values were determined by computer as best fitting their model of expected flow drop with respect to expected pressure jump. Of the three parameters, only resistance was consistent and in the expected range. Compliance values varied greatly. Inertance values were negative and therefore incorrect.
Shaw et al. (1983) were the most recent to reintroduce a perturbational device. Their device closely resembled the APD, but with a perforated metal plate sliding into the flowpath between the mouth and pneumotach rather than a screen open segment rotating beyond the pneumotach. Resistance was calculated from the magnitude of the flow perturbation with respect to the resistance of the perforated plate. Healthy subjects and COPD patients were tested. Resistance values were found to match plethysmographic resistance. Resistance measurements were found to vary with the number of holes in the perforated plate. Paradoxically, the healthy subjects used to demonstrate how the measurement changed with hole count all had higher resistances than the healthy subjects used to correlate perturbational resistance to plethysmographic resistance.
The literature is not consistent in its definition of perturbational resistance. Some authors treat the effective origin of the pressure driving respiration as the alveoli and assume that the perturbation does not extend through tissue. Therefore, perturbational resistance is airways resistance. They have reported generally lower resistance values. The other authors state that perturbation measures total respiratory resistance. They generally report higher resistance values.